The present invention generally relates to a method for forming hollow fibers.
Artificial gas exchange to and from biological fluids is widely performed in the clinical settings as well as in the laboratory for research purposes. In the clinical setting the use of blood oxygenators during extracorporeal life support (ECLS), is commonly used to replace or aid in the function of the lungs during heart surgery, or long-term cardiorespiratory support. Gas exchange is accomplished by creating a diffusion potential across a gas permeable membrane that drives a transfer of gas from a high partial pressure on one side, to a low partial pressure on the other side of the membrane, as provided by Fick""s law of diffusion:             m      .        g    =      K    ⁢          A      t        ⁢          (                        p          g          1                -                  p          g          2                    )      
where {dot over (m)}g is the rate of gas mass transfer across the membrane, K is a factor that is proportional to the solubility of the gas in the membrane and to the membrane diffusion coefficient, A is the membrane surface area, t is the membrane thickness and pg is the gas partial pressure on either side of the membrane. This equation suggests that for a given partial pressure difference across a membrane, more gas transfer is accomplished by increasing the membrane surface area, and by decreasing the membrane thickness.
Blood oxygenating devices that make use of gas transfer across a membrane come in a variety of styles. One design makes use of microporous membrane hollow fibers (MMHF). Microporous membrane hollow fibers are very small hollow tubes, with a typical outer diameter between 250 and 380 microns (xcexcm) and a typical wall thickness of about 50 microns (xcexcm). Multiple microporous membrane hollow fibers are typically wound into a fiber bundle, using a desired weaving pattern among the fibers. A typical fiber bundle constructed of microporous membrane hollow fibers is shown at 10 in FIG. 1. The individual fibers 12 forming the bundle 10 are illustrated with a greatly exaggerated diameter and wall thickness.
Blood oxygenating devices utilizing microporous membrane hollow fibers have become common for use during short-term cardiorespiratory support for procedures, such as routine bypass operations. In such a device, the ends 14 and 16 of the bundle 10 are each firmly potted into a potting material, which interconnects and seals to the ends of the fibers 12. A portion of the potting material is then sliced off so as to expose the hollow lumens 18 of each of the fibers 12. The potting material serves to interconnect and seal the outer surfaces of each of the fibers so that they are manifolded together at both ends. The potted bundle is then positioned in the oxygenator housing such that gas may be introduced into the lumens 18 while blood is passed over the outer surfaces of the fibers. Then, as oxygen flows inside the hollow fibers and blood flows over the outside of the fibers, the blood picks up oxygen and releases carbon dioxide across the microporous membrane, by diffusion.
In the fiber bundle, the walls of each fiber act as gas exchange membrane. Therefore, it is possible to compress a large membrane surface area into a relatively compact volume. In addition, as the blood flows outside the fibers, an increased convective mixing is achieved since the fibers downstream are within the wake or xe2x80x9ceddiesxe2x80x9d of those upstream. In this description, xe2x80x9cmembrane hollow fibersxe2x80x9d refers to hollow fibers where the walls of the hollow fibers act as membranes, and are typically thin to facilitate the transfer of mass and energy across the walls.
Microporous membrane hollow fibers are notorious for suffering from fowling and plasma leakage when they are used for extended periods of time: The blood plasma eventually leaks through the pores, thus compromising gas exchange, or rendering it completely ineffective.
Manufacturers of microporous membrane hollow fibers are incessantly seeking solutions to the plasma leakage problem, such as developing smaller pore size membranes that presumably have lower incidence of plasma leakage. Yet, no reports have been published showing improvement. Mitsubishi Rayon (Tokyo, Japan) introduced a multi-layered composite hollow fiber membrane (MHF) that contains a polyurethane interlayer sandwiched between two microporous polyethylene supporting layers. However, polyurethane has poor gas transfer properties, and fowling can still occur on the microporous side exposed to blood. There have been a number of other attempts to add dense coatings over microporous hollow fibers, yet none are available commercially. Notwithstanding the commercial availability of coated microporous membrane hollow fibers, the gas transfer through microporous membranes coated with silicone is reduced compared to the microporous membrane alone; gas must diffuse through the solid membrane in addition to the microporous membrane. Therefore, the tradeoff is reduced gas transfer.
Yet another potential problem associated with microporous membrane hollow fibers oxygenators is that if the gas side pressure becomes higher than the blood side, air can be readily transmitted through the micropores into the blood. Gas embolization may have fatal consequences if the gas bubbles are pumped into the patient. This can occur if the ports designed to vent the gas to atmosphere become occluded or if water condensation accumulates inside the lumen of the fibers, thus plugging the exhaust of oxygen. Consequently, gas side pressure must always be below the blood pressure to prevent gas embolization.
Because of the plasma leakage problem with microporous membrane hollow fibers oxygenators, spiral coil silicone membrane lungs (Medtronic Perfusion Systems, Brooklyn Park, Minn.), also known as Kolobow oxygenators, are used in long term applications because they do not have a propensity for plasma leakage. However, these solid membrane oxygenators require almost twice the surface area to achieve the same gas exchange as microporous hollow fibers. This is not because the membranes are not microporous, but because of the lack of convective mixing achievable over relatively xe2x80x9cflatxe2x80x9d membranes, compared to the mixing achievable over a bundle of thousands of hollow fibers. It should be noted that the oxygenated blood boundary layer, and not the membrane itself presents the major obstacle to oxygen diffusion to the blood.
A possible solution to the leakage problem with microporous membrane hollow fibers is to instead form hollow fibers out of a material that is not microporous, such as silicone. Gas diffusion can still occur across a silicone membrane, without the risk of gas embolization and plasma leakage. However, manufacture of blood oxygenators using silicone hollow fibers has not been commercially realistic.
Small silicone fibers can be extruded by polymer extruders in sizes comparable to the microporous polypropylene fibers. However, with the prior art there are two major barriers to the development of practical gas exchange devices. First, extrusion of solid silicone fibers is much more difficult and slower than extruding microporous polypropylene hollow fibers. Polypropylene is a thermoplastic polymer, whereas silicone is a cross-linked thermoset polymer. This means that polypropylene can be heated up, melted, and drawn-down to small diameters by pulling the extrudate as it comes out of the die, similar to making micropipettes with molten glass tubes. This allows for a significant reduction of fiber diameter from a manageable size die. Moreover, the polymer can be cooled quickly by water quenching once the fiber has been appropriately sized. Additional proprietary stretching processes are applied to render the fiber microporous.
Silicone, on the other hand, starts out as clay-like material that is extruded cold through the die (still as a clay) and then is heated to cure or cross-link the polymer, with much more limited drawdown compared to thermoplastics. Moreover, the clay like material has very little strength and is not as forgiving as molten plastic. As a result, the extrudate must be cross-linked or vulcanized quickly in order to control the tiny fiber. Thus, silicone extrusion is significantly slower than that of polypropylene to allow for polymer cure and subsequent handling such as winding in spools and fiber bundles. Further complicating the extrusion of tiny silicone fibers is the significant static buildup as the fiber cures through the oven. The static electricity makes it difficult to handle and wind the silicone fibers, especially if multiple fibers are extruded simultaneously.
The difficulty and the time-consuming process necessary to produce tiny silicone fibers are reflected in the product pricing. The price in 2002 for extruding one meter of silicone fiber (350 xcexcm OD, 250 xcexcm ID) was 33 cents, or $333 per kilometer (Specialty Manufacturing, Midland, Mich.). This compares to the microporous polypropylene hollow fiber price of $16 per kilometer (Celgard X30 240, Hoechst Celanese, Charlotte, N.C.). Thus the cost of the silicone fibers alone required for a device with 2.5 m2 of diffusion surface area could cost as much as $750.
Secondly, manufacturing oxygenators with silicone fibers is also significantly more difficult than with microporous membrane hollow fibers because silicone is elastic and flimsy. The winding of the fiber bundle becomes much more challenging and thus significantly slower. Moreover, potting the fiber bundle is not easy since the elastic silicone fibers tend to deform as the fiber bundle is subjected to large forces during centrifuging. Manufacturing of silicone fiber oxygenators is not impossible, but very difficult and therefore prohibitively expensive.
In view of the forgoing limitations and shortcomings of the prior art, as well as other disadvantages not specifically mentioned above, it should be apparent that there exists a need in the art for an improved gas permeable hollow fiber.
The present invention overcomes many of the shortcomings of the prior art by providing a method for producing a membrane hollow fiber which is not susceptible to plasma leakage and gas embolization. According to the present invention, a thin-walled microtube is formed by providing a continuous elongated member having an outer surface. The elongated member is at least partially formed of a water soluble material. A coating material is then provided, with the coating material being a silicone compound. The silicone structure is curable so as to form a substantially non-porous silicone. The outer surface of the elongated member is coated with the coating material so as to form a substantially uniform and continuous layer on the outer surface of the elongated member. The layer of coating material is then cured, and the elongated member is dissolved and purged from the layer of coating material. This leaves a micro tube formed of silicone. According to a further aspect of the present invention, the microtube formed according to the present invention may be assembled into a bundle and potted into a potting material at its ends prior to dissolving and purging the elongated member contained therein. After the fiber bundle is formed and potted, the elongated members in each of the microtubes may be dissolved and purged.